The present disclosure relates to medical imaging. More particularly, the present disclosure relates to direct detection x-ray imaging devices.
The most common and important application of fluoroscopic x-ray imaging today is in image guided cardiac therapeutic procedures where real-time (30 frame per second) image sequences are used to guide the interventional radiologist's mind and hand. One example is treatment of coronary artery disease (CAD) which results in thickening of the artery wall leading to a narrowing of the lumen and increased risk of thrombus formation. Fluoroscopy-guided catheterization is the method of choice for the investigation and treatment of CAD. The vessels are made visible by the injection of an iodinated radio-opaque contrast agent and images are obtained in real-time.
In 2010, it is estimated that there were over 3 million cardiac catheterization procedures performed in North America using x-ray fluoroscopy. The popularity of such procedures is caused by the fact, that they allow to replace open heart surgeries and thus less invasive. However, although these procedures are a great boon to patient care, they come at a price. First, the interventional cases tend to be longer than diagnostic procedures and can take 1-2 hours. With a typical fluoroscopic patient entrance exposure rate of 3 R/min (30 mGy/min skin dose), the skin dose from such procedures can be of the order of Gy and can approach the level where the patient is subject to somatic effects from a single procedure. In addition, the lifetime risk to the patient of a radiation induced cancer can be substantial, though difficult to calculate for an individual.
In short, there are many procedures currently in use which, for a single diagnostic or treatment session, can increase the probability of death from a subsequent radiation induced cancer by 1 in 200.
In addition to a risk to a patient, there are significant risks to interventional radiologists performing the procedures. While it is difficult to precisely quantify risk of malignancy in physicians using fluoroscopy, the general consensus is that there is an increased risk. Hence, x-ray exposure to patient and medical personnel during fluoroscopic procedures has to be reduced, without sacrificing the image quality.
Flat panel x-ray detectors based on active matrix flat panel imagers (AMFPI) are used in state-of-the-art fluoroscopic systems. Currently, flat panel fluoroscopic systems employ an indirect conversion scheme, in which a Csl scintillator first converts x-ray quanta into optical photons, which in turn diffuse through a phosphor and then are converted back to electrons by an array of photodiodes. This indirect and multi-stage conversion process reduces the conversion gain, while the resolution of the detector degrades as a result of the isotropic light spread that occurs even when the scintillators are structured. The aforementioned problems associated with indirect detection can be addressed by the use of direct conversion detectors, where a photoconductive layer is deposited directly on an AMFPI and acts as x-ray-to-charge transducer. X-rays are absorbed in a photoconductor that directly creates electron hole pairs, which are separated and moved by an electric field and thus there is no significant loss of resolution. By reducing the number of stages, the conversion process can be up to ten times more efficient than for scintillator, making it more efficient at the lowest exposure rates.
Direct detection requires a photoconductive having a distinct set of properties. Four important photoconductor properties for direct x-ray detection are: (1) high conversion gain; (2) good photoconductive properties; (3) high absorption efficiency and (4) compatibility with large area detector technology. Currently, the only commercially viable x-ray photoconductor in direct conversion x-ray detectors is a-Se. Unfortunately, a-Se is a low-Z (atomic number) material and thus has adequate absorption only at low x-ray energies and the high exposures (i.e. exposures suitable for digital mammography), while at the lowest fluoroscopic doses, a-Se direct conversion FPDs have similar conversion efficiency as Csl indirect detectors. In order to achieve suitable imaging performance for low-dose fluoroscopic procedures, a-Se has to be replaced with a high-Z material that has high absorption and also possesses lower electron-hole pair creation energy, and therefore a higher conversion gain.
Polycrystalline lead oxide (PbO) satisfies all criteria since:
(1) It has a theoretically predicted high conversion gain;
(2) Its appropriate photoconductive properties have been proven by applications in Plumbicon video pick-up tubes; and
(3) It has a higher X-ray detection quantum efficiency due to its high Z.
Hence, a flat panel direct conversion detector based on PbO technology would appear to possess the features needed to meet the requirements of fluoroscopic cardiac interventional procedures. Polycrystalline PbO was previously shown to have many of the requirements for an effective photoconductive material and additionally it has been previous used in small area imaging systems (Plumbicons). This indicates an adequate temporal response when used in thin layers, while a thicker layer and larger coated area are needed for medical imaging applications.
In 2005 Simon et al. demonstrated a complete large area flat panel imager, indicating that the PbO deposition process is compatible with a-Si electronics and allows large detector area coating (M. Simon, R. A. Ford, A. R. Franklin, S. P. Grabowski, B. Menser, A. Nascetti, M. Overdick, M. J. Powell, D. U. Wiechert, “Analysis of Lead Oxide (PbO) Layers for Direct Conversion X-Ray Detection”, IEEE vol. 52, 2037(2005)). Unfortunately, the PbO layers manufactured using conventional deposition techniques were very porous. The film exhibited a rough surface morphology and composed of randomly oriented platelets several micron in diameter and a few hundred nanometers thick. The density of the grown layers was much lower than that of a crystalline material (up to 50% of single crystal density), which significantly decreases the X-ray attenuation of the grown film.
Furthermore, the grown PbO films are known to consist of two different crystallographic phases of PbO: the seeding layer, several microns thick, is formed by the yellow orthorhombic PbO with band gap of 2.7 eV, while the bulk of the layer grows as a red tetragonal lead oxide with band gap of 1.9 eV. The presence of an orthorhombic phase diminishes detector performance, and leads to the requirement of post-growth treatment of the PbO layer. In addition, the deposited films are unstable in air and known to degrade in the ambient environment.
In addition, PbO photoconductive layers have not yet shown adequate temporal behavior for fluoroscopic applications. The films are reported to exhibit significant image lag (the percentage of residual signal present in a subsequent frame), which precludes their use in real time imaging (i.e. dynamic imaging used in fluoroscopy) and restricts their application to static imaging only (radiology).
It therefore follows that the full potential of PbO remains unexploited in view of the aforementioned technical problems and limitations.